Grating-based phase contrast imaging (PCI) is a relatively new imaging method by which, instead of an X-ray absorption image, measurement data is recorded which makes it possible to obtain in parallel both an absorption-based X-ray image, a differential phase image and a dark field image. From the data in the two additional images, further information can be obtained which can, for example, be used in a clinical diagnosis.
A requirement for grating-based phase contrast imaging is local coherence of the X-ray radiation, at least in one direction.
A further technological bottleneck is that the interference pattern which is to be recorded normally has a smaller periodic resolution than that which can be resolved by standard X-ray image detectors. The standard solution to this is to use a so-called phase-stepping technique, which will be explained in more detail below by reference to FIG. 2. With this, an analyzer grating G2 with the periodicity of the undisturbed interference pattern is set up in front of the X-ray image detector, and is displaced relative to the object and the X-ray image detector, where the magnitude of the displacement is less than one period of the grating. From three or more one-shot recordings, the interference pattern can then be reconstructed.
For differential phase contrast imaging (PCI), it is usual to insert three gratings in the beam path from the source of the X-ray beam. The article “Phase retrieval and differential phase-contrast imaging with low-brilliance X-ray sources” by Franz Pfeiffer et al., which appeared in Nature Physics 2 (2006), pages 258 to 261, describes an example of this type of PCI, which will be explained in more detail below.
The wave nature of particles such as X-ray quanta permits phenomena like diffraction and reflection to be described with the help of the complex refractive indexn=1−δ+iβ. 
Here, the imaginary component β describes the absorption, on which is based today's clinical X-ray imaging, such as for example computed tomography, angiography, radiography, fluoroscopy or mammography, and the real component δ describes the phase shift which is used in differential phase imaging.
From DE 10 2010 018 715 A1, an X-ray radiography system is known in which, for the purpose of high-quality X-ray imaging, for the purpose of phase contrast imaging of an object under investigation use is made of an X-ray radiography system which has at least one X-ray emitter with a plurality of field emission X-ray sources for emitting coherent X-ray radiation, an X-ray image detector, a diffraction grating G1 arranged between the object under investigation and the X-ray image detector, and another grating G2 which is arranged between the diffraction grating G1 and the X-ray image detector.
An X-ray recording system, with which differential phase contrast imaging of the type mentioned in the introduction can be carried out, is known for example from U.S. Pat. No. 7,500,784 B2, and this will be explained by reference to FIG. 1.
FIG. 1 shows the typical main features of an X-ray radiography system for an intervention suite with a C-arm 2, held by a stand 1 in the form of a six-axis industrial or articulated arm robot, attached to the ends of which are an X-ray radiation source, for example an X-ray emitter 3 with X-ray tubes and collimator, and an X-ray image detector 4 as the image recording unit.
Using, for example, the articulated arm robot known from U.S. Pat. No. 7,500,784 B2, which preferably has six axes of rotation and hence six degrees of freedom, the C-arm 2 can be moved to any desired position in space, for example by being turned about a center of rotation between the X-ray emitter 3 and the X-ray image detector 4. The inventive angiographic X-ray system 1 to 4 can, in particular, be rotated about centers of rotation and axes of rotation in the plane of the C-arm of the X-ray image detector 4, preferably about axes of rotation which pass through the mid-point of the X-ray image detector 4 and the mid-point of the X-ray image detector 4.
The familiar articulated arm robot has a basic frame which, for example, has a fixed mounting on the floor. To this is attached a carousel that can rotate about a first axis of rotation. Attached to the carousel so that it can pivot about a second axis of rotation is a robot swing arm, to which is affixed a robot arm which can rotate about a third axis of rotation. At the end of the robot arm is attached a robot hand which can rotate about a fourth axis of rotation. The robot hand has a fixing element for the C-arm 2, which can pivot about a fifth axis of rotation and can rotate about a sixth axis of rotation oriented at right angles to the fifth.
The implementation of the X-ray diagnostic facility is not dependent on an industrial robot. It is also possible to make use of the usual C-arm or mammography devices.
The X-ray image detector 4 can be a rectangular or square flat semiconductor detector, preferably manufactured from amorphous silicon (a-Si) or selenium (a-Se). However, use could also be made of integrating and possibly counting CMOS detectors.
In the beam path from the X-ray emitter 3 there is, on a table plate 5 of a patient positioning table, a patient 6 who is to be investigated as the object under investigation. Connected to the X-ray diagnostic facility is a system control unit 7 with an imaging system 8, which receives and processes the image signals from the X-ray image detector 4 (examples of the operating elements are not shown). The X-ray images can then be inspected on displays on a suspended monitor 9. The suspended monitor 9 can be held by means of a ceiling-mounted, longitudinally movable, pivotable, rotatable and height-adjustable carrier system 10 with a cross-arm and lowerable carrier arm.
Instead of the X-ray system illustrated by way of example in FIG. 1, with its stand 1 in the form of a six-axis industrial or articulated arm robot, the angiographic X-ray system could also have a normal ceiling or floor mounted holder for the C-arm 2.
Instead of the C-arm 2 illustrated by way of example, the angiographic X-ray system could also have separate ceiling and/or floor mounted holders for the X-ray emitter 3 and the X-ray image detector 4, which are, for example, rigidly coupled to each other electronically.
In today's known arrangements for clinical phase contrast imaging, use is made of conventional X-ray tubes, X-ray image detectors which are today available, such as for example those described by Martin Spahn in “Flachbilddetektoren in der Röntgendiagnostik” [Planar image detectors in X-ray diagnostics], Der Radiologe, Volume 43 (5-2003), pages 340 to 350, and three gratings, G0, G1 and G2, as will be explained in more detail by reference to FIG. 2, which shows schematically a structure for a Talbot-Lau interferometer for differential phase contrast imaging, with extended tube focus, gratings G0, G1 and G2 and a pixelated X-ray image detector.
For the purpose of producing coherent radiation, the X-ray beams 12 emerging from a tube focus 11 of the non-coherent X-ray emitter 3 pass through an absorption grating 13 (G0) which effects local coherence, and through an object under investigation 14, for example the patient 6. The object under investigation 14 deflects the wave front of the X-ray beam 12 by phase shifting in a way that is made clear by the normal 15 to the wave front when there is no phase shift, i.e. with no object, and the normal 16 to the wave front with phase shifting. Following this, the phase-shifted wave front passes through a diffraction or phase grating 17 (G1), which has a grating constant matched to the mean energy of the X-ray spectrum, for the purpose of producing interference lines (Talbot effect), and then in turn through an absorptive analyzer grating 18 (G2) for reading out the interference pattern produced. The grating constant of the analyzer grating 18 is matched to that of the phase grating 17 and to the remaining geometry of the arrangement. The analyzer grating 18 is, for example, arranged at the first or nth Talbot interval. By so-called “phase-stepping”, described below, together with the analyzer grating 18 (G2), it is possible to detect relevant items of data from the interference pattern, using the X-ray image detector 4.
If the tube focus 11 of the X-ray radiation source is sufficiently small, and the radiation power generated is nevertheless sufficiently large, it may be possible to forgo the first grating G0, the absorption grating 13, as would be the case for example if a plurality of field emission X-ray sources are provided as the X-ray emitter 3, as is known from DE 10 2010 018 715 A1 which has been described.
For each pixel of the X-ray image detector 4, the image data is now determined in that, by the phase stepping 19, which is indicated by an arrow, the analyzer grating 18 (G2) is moved in several steps (k=1, K, where for example K=3 to 8) by an appropriate fraction of the grating constant, perpendicularly to the direction of radiation of the X-ray beams 12 and laterally relative to the arrangement of the grating structure, and the signal Sk which arises for this configuration during the recording is measured in the pixel of the X-ray image detector 4, and by this means the interference pattern which arose is sampled. For each pixel, the parameters of a function (e.g. a sine function) which defines the modulation is then determined by a suitable fitting method, by a matching or compensation method applied to the signals Sk measured in this way. The visibility, i.e. the normalized difference between the maximal and minimal signals (or more precisely: the amplitude normalized to the mean signal), is here a measure for characterizing the quality of a Talbot-Lau interferometer. It is defined as the contrast of the sampled modulation
  V  =                              I          max                -                  I          min                                      I          max                +                  I          min                      =                  A                  I          _                    .      
Further, in this equation A represents the amplitude and Ī the mean intensity. The visibility can take on values between zero and one, because all the variables are positive and Imax>Imin. In a real interferometer it is also the case that Imin>0, so that the value of V meaningfully extends across the whole range. Minimal intensities greater than zero, and all non-ideal characteristics and defects of the interferometer, lead to a reduction in the visibility. A third item of data which can be defined by means of the visibility and is generated by this type of measurement is called the dark field. The dark field gives the ratio of the visibilities of measurements with an object and those without an object.
  D  =                    V        obj                    V        ref              =                                        A            obj                    ·                                    I              _                        ref                                                A            ref                    ·                                    I              _                        obj                              .      
From a comparison of certain derived quantities from the functions fitted for each pixel, once with and once without an object (or patient), it is then possible to produce three different images:
(i) an absorption image,
(ii) a differential phase contrast image (DPC) and
(iii) a dark-field image.
In the case of the known PCI imaging methods, use is currently made of either a microfocus X-ray tube, which of itself meets the required coherence conditions, or alternatively an absorption grating with the designation G0, which splits up the array of X-ray beams, output from the anode of the X-ray tube, into lines of X-rays. Each of these lines taken on its own satisfies the coherence condition and is so positioned by its spacing from its neighboring lines that in the plane of the detector the interference images overlay each other constructively in conformity with the method according to Lau.
The disadvantage of this method is that a large proportion of the X-ray radiation generated is absorbed in the absorption grating G0, because the ratio of the opening to absorptive material is ≦1.